Laser Based Metal Deposition LBMD of Antimicrobials to Implant Surfaces

ABSTRACT

A method is provided for depositing a hard wear resistant surface onto a porous or non-porous base material of a medical implant. The wear resistant surface of the medical implant device may be formed by a Laser Based Metal Deposition (LBMD) method such as Laser Engineered Net Shaping (LENS). The wear resistant surface may include a blend of multiple different biocompatible materials. Further, functionally graded layers of biocompatible materials may be used to form the wear resistant surface. Usage of a porous material for the base may promote bone ingrowth to allow the implant to fuse strongly with the bone of a host patient. The hard wear resistant surface provides device longevity, particularly when applied to bearing surfaces such as artificial joint bearing surfaces or a dental implant bearing surfaces. An antimicrobial material such as silver may be deposited in combination with a metal to form an antimicrobial surface deposit.

CROSS-REFERENCE TO RELATED DOCUMENTS

This application is a divisional of:

Pending U.S. patent application Ser. No. 11/624,041, filed Jan. 17, 2007by Daniel F. Justin, et al. for LASER BASED METAL DEPOSITION (LBMD) OFANTIMICROBIALS TO IMPLANT SURFACES, (Attorney's Docket No. MLI-58NPROV), which is incorporated by reference, and is a non-provisional of:

Prior U.S. Provisional Patent Application No. 60/811,934, filed Jun. 7,2006 by Brent E. Stucker, et al. for INCORPORATION OF SILVER INTOMATERIALS DEPOSITED BY LBMD, (Attorney's Docket No. MLI-58 PROV), whichis incorporated herein by reference.

The following are also incorporated herein by reference:

Issued U.S. patent application Ser. No. 11/432,426, filed May 10, 2006by Daniel F. Justin et al. for LASER BASED METAL DEPOSITION (LBMD) OFIMPLANT STRUCTURES (Attorney's Docket No. MLI-05 CIP) issued as U.S.Pat. No. 7,666,522 on Feb. 23, 2010;

Issued U.S. patent application Ser. No. 11/253,850 filed Oct. 18, 2005by Daniel F. Justin et al. for LASER BASED METRAL DEPOSITION OF IMPLANTSTRUCTURES, (Attorney's Docket No. MLI-05 CON) issued as U.S. Pat. No.7,632,575 on Dec. 15, 2009;

Issued U.S. patent application Ser. No. 10/811,038, filed Mar. 26, 2004by Daniel F. Justin et al. for LASER ENGINEERED NET SHAPING OF IMPLANTSTRUCTURES, (Attorney's Docket No. MLI-05 NPROV) issued as U.S. Pat. No.7,001,672 on Feb. 21, 2006; and

Prior U.S. Provisional Patent Application No. 60/527,118, filed Dec. 3,2003 by Daniel F. Justin et al. for LASER ENGINEERED NET SHAPING OFIMPLANT STRUCTURES, (Attorney's Docket No. MLI-05).

FIELD OF THE INVENTION

The present invention relates to the formation of biocompatiblematerials onto a medical implant device, and more particularly to theuse of laser based metal deposition of biocompatible materials onto baseimplant material structures.

BACKGROUND OF THE INVENTION

The advancement of enhanced materials for the use of medical implants,such as joint prostheses has immensely improved the quality of life formany people over the past century. Devices such as artificial hips,knees, shoulders and other devices have allowed people who wouldotherwise have suffered from chronic pain and physical limitation tolive active, comfortable lives. The development of such devices hasconfronted scientists and engineers with many technical challenges, suchas in the area of materials science engineering wherein to achieveoptimal implant performance various biocompatible materials withdifferent physical and mechanical properties are bonded to each other.

Materials used for such devices must not only be non-corrosive, but mustalso be sufficiently resilient (having high tensile and compressivestrength), and hard (having sufficient wear resistance). Since a devicesuch as an artificial joint must undergo a great number of cycles ofwear during the lifetime of the host patient, such devices must alsopossess great fatigue properties.

Some medical implant devices such as artificial joints must bond in someway with the patient's natural bone. Early devices employed bondingpolymers, commonly referred to as bone cement to bond the implantrigidly to the anatomic structure of bone. However, more recently suchdevices have been constructed of porous materials such as porousTitanium (Ti) and porous Tantalum (Ta). The bone of the host patientgrows into the porous material creating a strong permanent mechanicalbond without the use of bone cements. Consequently, such implants aremore reliable and durable in the long term than those relying on bonecement for fixation.

Such implant devices are typically manufactured from a wrought alloy,forged alloy or a powder metal injection molded process. While thisproduces an implant device with bulk properties that are optimized forcertain overall design criteria such as biocompatibility strength andmodulus of elasticity, these properties may not be optimized forproperty requirements specific to certain portions of the implant, suchas wear or bone ingrowth characteristics.

For instance, while the use of porous materials such as porous Tiprovides crucial and beneficial bonding properties, such materials maynot have optimal properties in other areas. For example, porousmaterials may not be as hard as some other biocompatible materials andtherefore may not have acceptable wear properties. However, because ofthe overriding importance of strong permanent bonding with the hostpatient bone, such porous materials have continued to be used in spiteof less than optimal wear properties.

In order to enhance the wear properties of a device such as anartificial joint, prior art devices have been constructed in more thanone piece. A first portion of the joint implant, that which will bond tothe bone, has typically been constructed of a porous material such asporous titanium, and a second piece, such as the bearing surface of thejoint has been constructed of a much harder, more wear resistantmaterial such as alloys of cobalt and chrome (Co—Cr). The first andsecond pieces are then bonded together in an attempt to obtain thebenefits of both materials. One challenge to using such a technique isthat of achieving a sufficiently strong, permanent bond between thefirst and second portions, without the use of adhesives that may bebiologically incompatible or may fail under the stresses imposed by thebody of the patient. Attempting to weld such materials together cancause the non-porous material to flow into the porous material,destroying the porosity of the porous material and degrading the abilityof the device to bond with the patient's bone. In addition, suchmaterials, being dissimilar metals, often experience galvanic corrosionwhen bonded together in such a manner.

Therefore, there remains need for a device (and method for making thesame) such as an artificial joint which can take advantage of theproperties of a first material, such as the porosity of porous Ta or Ti,and also take advantage of the properties of a second material, such asthe hardness of a material like Co—Cr, for use in a bearing environmentsuch as a ball or socket of a joint. Such a device would preferably notexhibit any delamination between the two materials and would notexperience any galvanic corrosion. Such a device would also preferablynot diminish the porosity of the porous material due to the flow of theother material thereinto.

SUMMARY OF INVENTION

The present invention provides a method for constructing a medicalimplant such as a hip prosthesis, having a bulk portion constructed of aporous material which can fuse with a host patient's bone structure, andwhich also has a hard, wear resistant material only at portions of thedevice where such properties are desired. According to the invention, alaser based metal deposition (LBMD) layer of relatively dense hardmaterial, can be applied to a porous material base structure or anon-porous material base structure.

The relatively hard, wear resistant biocompatible material can be forexample an alloy of cobalt and chrome alloy, whereas the porous materialcould be a biocompatible material conducive to bony tissue ingrowth whenformed in a porous structure such as porous Titanium, Ti6Al4V, Ti6Al4VELI, Titanium-Nickel alloys, Tantalum, Tantalum alloys, and porousstructures made from other materials that have an exposed surface madefrom biocompatible materials. Some applications may call for anon-porous base material, in which case suitable base materials includeTitanium, Ti6Al4V, Ti6Al4V ELI, Titanium-Nickel alloys, Niobium,Zirconium, Tantalum, stainless steels and alloys related to thesemetals. In these applications, the relatively hard wear resistantbiocompatible material can comprise of materials such as carbides,oxides, nitrides or other compositions that are harder or more wearresistant than the non-porous biocompatible base materials.

According to the LBMD material application of the present invention, theapplied material can be applied as, for example, powdered metal, as awire or as a foil. The applied material is then melted by a high-energylaser immediately upon or soon after application. The use of a laser toheat the applied material advantageously allows the heating to be verylocalized, thereby minimizing any adverse effects of such heat on theunderlying material. However, in this invention, other sources of energysuch as plasma and electron beam energy could also be focused andlocalized to allow energy based metal deposition of material throughsimilar processes. Thus, in this invention laser energy is the preferredenergy source. However it is understood that the term Laser Based MetalDeposition (LBMD) is meant to include any other energy source that canprovide sufficient energy locally to melt and deposit the additivematerial, but minimize any adverse effects of such heat on theunderlying base material.

In addition, the extremely localized heating of the laser, or any otherlocalized energy source, in conjunction with the heat sinking propertiesof the underlying material leads to very rapid subsequent cooling,resulting in a beneficial small grain structure as well as allows theaddition of carbon interspersions when conducted in a carbon-richenvironment or with powdered or alloyed carbon added to the depositionmaterial, both of which provide increased hardness to the depositedmaterial.

Furthermore, since the LBMD deposited material is heated and cooled soquickly and locally, the applied material tends not to flow excessivelyinto the porous or underlying non-porous base material, therebymaintaining the desirable porous properties of the porous bulk portionof the device or preferred mechanical and physical properties of thenon-porous base material. This also creates a relatively small bondingzone between the base material and the LBMD deposited material. Thisallows for a thin layer of LBMD deposited material to be deposited ontothe base material. Because this layer of deposited material is thin,implants can be fabricated that are optimized in size to limit theamount of bone that must be removed to facilitate the bulk of theimplant. For example, a 5 millimeter thick sheet-like implant with a 3millimeter thick porous bone ingrowth underside, a 0.5 millimeterbonding zone, and 1.5 millimeter bearing surface made from a first layerof Titanium and a second layer of Cobalt-Chrome can be placed as bearingpads on the proximal tibial plateau as a tibial hemiplasty implant inthe knee. This construct of the 5 millimeter thick implant issignificantly bone conserving compared to traditional 9 millimeter to 20millimeter thick tibial implants that are currently used to resurfacethe proximal tibia of the knee.

In another aspect of the invention, a relatively hard material such asCo—Cr can be applied to the surface of a porous base such as porousTantalum, and the Co—Cr surface used to bond to a Co—Cr bulk portion ofthe device. This overcomes the problems that have previously beenexperienced, when trying to bond a material such as Co—Cr to anothermaterial such as porous Tantalum. A corrosion barrier, such as a layerof Ti may be provided between the porous Tantalum and the Co—Cr.

Another aspect of the invention it that the invention allows for moreeconomical fabrication of implantable devices. For example, themanufacturing process can be optimized utilizing aspects of thisinvention by making the base structure from a readily manufacturablematerial and depositing a second material on to the base material usingLBMD. Biocompatible materials with inherently greater hardness and moredifficult machinability characteristics, such as Titanium Carbide, canbe selectively deposited onto more easily machinable materials such as awrought Titanium alloys. This allows for the majority of the implantmade from the more easily machinable wrought alloy to be fabricated intothe base shape of the implant. Then, only the articulating bearingsurfaces of the implant need to be fabricated from the LBMD depositedharder material such as titanium carbide (TiC). In this example,titanium carbide is used as the harder LBMD material. However, anyarticulating bearing material may be deposited on any more easilymachined base material as long as the specific metallurgic chemistrybetween the two materials allows for bonding between the two materialswithout the excessive addition of tertiary phases of interstitialmaterials that prevent adequate bonding.

The present invention can also allow use of a blended composite coatingin order to enhance the coating performance. This approach allows one totailor the coating properties as desired by choosing a suitable matrixand reinforcing phases. The matrix material, such as Titanium, addressesthe considerations of metallurgical compatibility, metallurgical bondingand coating brittleness, while the reinforcing second phase material,such as Titanium Carbide, takes care of the hardness and wear resistanceissues. In addition, the invention incorporates the principles offunctionally graded materials, either to achieve desired propertygradation across the coating thickness or to facilitate satisfactorycoating deposition or both. Functionally graded materials (FGMs) arecomposite materials where the composition or the microstructure arelocally varied so that a certain variation of the local materialproperties is achieved. This is achieved by closely controlling theamount and distribution of second phase particles across the coatingthickness. The current approach helps achieve superior coatingproperties as well as minimize cracking problems during coatingdeposition.

The invention takes advantage of the capabilities of advanced directmetal deposition techniques, specifically, deposition of a compositecoating using a laser based metal deposition process. Some of theadvantages of this method include: i) the coating can be metallurgicallybonded to the substrate, ii) minimal heat input and very narrow heataffected zones, iii) spatial variation in coating microstructure andproperties is possible, iv) coating microstructure can be closelycontrolled, v) reduced residual stresses and distortion problems, andvi) precise control over coating application area. These capabilitiesare not available with conventional coating processes.

The invention produces coatings which are significantly thicker thanmost existing coating techniques. Existing techniques such as ionimplantation or surface treatments of Titanium nitride or Titanium oxideproduce surface treatments with thicknesses measured at the molecular ornanoparticle level, whereas the invention produces a coating thickenough to create a functional bearing surface. Compared to therelatively thin surface treatments described above, the LBMD thicknessgenerally ranges from a fraction of a millimeter to multiple millimetersin thickness. The thickness of the coating is important, in thatconcentrated stresses at the surface of a thin coating result inconcentrated stresses at the coating/substrate interface, which can leadto delamination. In addition, a certain amount of wear is expected whentwo materials are articulating against each other. In the case of therelatively thin existing surface treatment techniques, these materialswear away from the surface causing the softer underlying base materialto be exposed. In contrast, load is transmitted through thickercoatings, the concentrated load at the surface passes through a largervolume of coating, and becomes distributed over a larger area as thestress reaches the coating/substrate interface, resulting in a lowerconcentration of stresses at the location of greatest failure potential.

The present invention provides a manufacturing method for producing animplant made from traditional or novel implant metals with layers ofmaterial having differing densities and structures.

The present invention provides a surface material deposition processthat allows for a gradient of materials with varying selectiveproperties to be deposited on the bulk implant material. After the basestructure is formed, additional material is added to the base structureusing the LBMD process.

The implant is formed in the approximate final shape from a common ornovel orthopedic alloy such as Co—Cr alloys, titanium alloys, stainlesssteel alloys, or base pure metal such as tantalum, titanium or platinum.Because the basic structure of the implant is formed by conventionalmanufacturing means out of implant grade materials, the majority of thecost of the manufacturing is similar to existing implants.

Applicable implant shapes that can benefit from LBMD deposition ofharder materials onto the base material include knee, shoulder, hip,finger, spine, top, foot, elbow, wrist, dental, jaw, and ankleprosthesis, just to name a few.

Besides improving bearing properties of implants, the LBMD process canbe used to increase the bone ingrowth properties of implant surfaces.This can be done by either depositing a hard material onto a porous basematerial or depositing a porous material onto a hard material.

In the case of adding a hard material to a base material, a monoblock ofa porous structure of an implant material is the base material. Aclosely packed fine grain structure of an implant material is then addedto the base material by LBMD methods. The closely packed grain structurewould result in improved wear properties.

The majority of the bulk of the implant can be manufactured byconventional methods. The hardened surface may then be added by LBMDdeposition. Unlike structures that are completely made by methods suchas LBMD, this method would allow the majority of the structure to bebuilt by conventional methods with only thin layers of hard materialadded to the structure. Accordingly, cost savings can be achieved.

LBMD allows for a highly focused laser beam of energy to melt a verysmall amount of powder over a short period of time. Because the largebulk material acts as a heat sink, this process results in a rapidlycooled LBMD deposited material. Rapid cooling of materials such asmetals results in a finer grain structure, which results in increasedhardness. In addition, in a carbon rich environment, carbides formresulting in an even harder material. Since the hardness of a materialis typically directly related to wear resistance, materials having highhardness become very attractive for use on bearing surfaces such asthose on knee, hip, wrist and elbow joints as well as myriad otherimplant devices.

Using the material deposition process of the present invention, likematerials can be deposited onto like materials such as Co—Cr alloys LBMDdeposited on Co—Cr wrought materials. However, dissimilar materials mayalso be deposited, such as titanium alloys deposited on Co—Cr alloys, orCo—Cr alloys can be deposited on titanium and its alloys.

In selected embodiments of the present invention, materials havingantimicrobial properties may be deposited on a metal base structure ofan implant through the use of LBMD, with or without LENS techniques.Such materials may include, but are not limited to elemental silver,gold, platinum, palladium, iridium, copper, tin, antimony, bismuth,zinc, salts thereof, and intermetallics thereof. The antimicrobialmaterial may be deposited along with other materials, such as a metalselected from cobalt-chrome, tantalum, titanium, platinum, zirconium,niobium, stainless steel, and alloys thereof. The deposit including theantimicrobial material may be formed as a single layer or as multiplelayers, with or without functional gradients provided by variation inthe proportions of the component materials deposited.

The systems and methods of the present invention may be used to providea number of different types of implant surfaces, including boneapposition surfaces, articular surfaces, and surfaces that do not abutbone or articulate. Bone apposition surfaces may be deposited with arelatively high porosity to enhance bone in-growth. Articular surfacesmay be formed with a higher hardness to provide enhanced wearresistance. A wide variety of implant surfaces, including those that donot articulate or contact bone, may desirably be formed withantimicrobial properties to reduce the probability of infection of thetissues surrounding the implantation site.

Other aspects and advantages of the present invention will becomeapparent from the following detailed description, which, when taken inconjunction with the drawings, illustrate by way of example theprinciples of the invention.

BRIEF DESCRIPTION OF THE DRAWINGS

For a fuller understanding of the nature and advantages of thisinvention, as well as the preferred mode of use, reference should bemade to the following detailed description read in conjunction with theaccompanying drawings.

FIG. 1 shows an example of the present invention employed in a hipprosthesis;

FIG. 2 is a view taken from circle 2 of FIG. 1, showing the a crosssection of the surface of the hip prosthesis of FIG. 1;

FIG. 3A illustrates the deposition of a first material using laser basedmetal deposition (LBMD);

FIG. 3B illustrates the deposition of a second material on the firstmaterial of FIG. 3A using laser based metal deposition (LBMD);

FIG. 3C is a micrograph at 5× magnification that shows three layers ofCo—Cr alloy deposited by the LBMD process on a bulk material of wroughtCo—Cr;

FIG. 3D is a micrograph at 5× magnification of nine layers of Co—Cralloy deposited by LBMD on a bulk material of wrought Co—Cr;

FIG. 3E is a micrograph at 50× magnification showing the bulk wroughtCo—Cr alloy;

FIG. 3F is a micrograph at 50× magnification showing the LBMD depositedCo—Cr alloy, particularly showing the finer grain structure associatedwith a rapidly cooled LBMD deposited material;

FIG. 4 illustrates an alternate application of the present invention;

FIG. 5 shows various implants that could have improved bone ingrowths orbearing properties if processed by LBMD;

FIG. 6 is a partial cross sectional view of the toe implant of FIG. 5taken along line 6-6 of FIG. 5;

FIG. 7 is a partial cross sectional view of the dental implant of FIG. 5taken along line 7-7 of FIG. 5;

FIG. 8 is a partial cross sectional view of one articulating implant ofFIG. 5 taken along line 8-8 of FIG. 5.;

FIG. 9 is a partial cross sectional view of the thumb implant 508 ofFIG. 5 taken along line 9-9 of FIG. 5;

FIG. 10 is an exploded view the knee implant of FIG. 5 and a multi-layerstructure coupling thereto;

FIG. 11 illustrates the deposition of multiple layers of a powdermaterial blend using LBMD;

FIG. 12 illustrates the deposition of a single layer of the powdermaterial blend using LBMD;

FIG. 13 illustrates the deposition of multiple layers of the firstmaterial using LBMD;

FIG. 14 illustrates the deposition of functionally graded layers of thefirst material by one nozzle and the second material by another nozzleusing LBMD;

FIG. 15 illustrates the deposition of multiple layers of a blend of anantimicrobial material and a first metal using LBMD;

FIG. 16 illustrates the deposition of single layer of the blend of theantimicrobial material and the first metal using LBMD;

FIG. 17 illustrates the deposition of functionally graded layers of theantimicrobial material and the first metal using LBMD; and

FIG. 18 is perspective view of the knee implant of FIG. 10 showingzone-specific deposits of bearing surface material, porous bone ingrowthmaterial and antimicrobial material.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

The following description is the best embodiment presently contemplatedfor carrying out this invention. This description is made for thepurpose of illustrating the general principles of this invention and isnot meant to limit the inventive concepts claimed herein.

With reference to FIG. 1, a preferred embodiment of the presentinvention will be described in terms of a hip prosthesis (hip) 100 forimplanting in the body of a patient. However, this is only by way ofexample, and it should be understood the present invention can practicedon many other medically implanted devices, including without limitation,knee, shoulder and elbow prostheses, as well as many other devices. NoteFIG. 5, discussed below.

The hip prostheses 100 must be constructed completely of biocompatiblematerials in order to ensure acceptance of the prostheses by thepatient's body. A biocompatible material is one that will not cause anadverse reaction with a host patient, and that will not corrode whenexposed to human tissue and fluids within the host patient. The hip 100includes a base portion 102, which may include a shank 104 and a ball106, and that is constructed predominantly or completely of a porousmaterial such as porous Ti or Ta (or alloys thereof). Constructing theshank 104 of a porous material such as Ti or Ta advantageously promotesbone growth into the porous material and strong fusion therewith. Thisprovides a strong, permanent, resilient bond with the bone of the hostpatient without the need for adhesives. As discussed above, the use ofadhesives to bond the hip 100 to the bone of the host patient would notonly provide a somewhat unreliable bond, but could also lead to adversereactions with the host patient.

As also mentioned above, the base 102 is constructed either completelyor predominantly of a porous material, such as a porous matrix of Ta orTa alloy, Ti or Ti alloy, for example Ti-6Al-4V, Ti—Ni, Ti6Al4V ELI,Titanium-Nickel alloys, and porous structures made from other materialsthat have an exposed surface made from biocompatible materials. The base102 can be formed by methods such as casting, machining or forging.

A preferred material for the base 102 is porous tantalum. One suchporous tantalum is sold under the brand name HEDROCEL® by IMPLEX®Corporation, 80 Commerce Drive, Allendale, N.J. 07401.

The preferred porous tantalum material such as HEDROCEL® has an opencell, tantalum metal structure that has the appearance of cancellousbone, and that can be formed or machined into complex shapes. It isdistinguished from current porous materials by its uniformity andstructural continuity as well as by its strength, toughness, andresistance to fatigue failure.

The tantalum metal structure consists of interconnecting pores,resulting in a construct that is >60% porous, and ideally >75% porous.In addition, the tantalum material preferably has flexural modulusproperties that are similar to those of human bone. For articulatingjoint replacement devices, compression molded polyethylene can beinfused into the tantalum structure, creating a bond as strong as thepolyethylene itself. In addition, the titanium structure can befabricated into products without the need for solid metal support.

The preferred porous tantalum metal (e.g., HEDROCEL®) has a similarcellular geometric appearance to bone graft, and also offers manybeneficial attributes. The porous structure is preferably a uniform andhomogeneous biomaterial, having load carrying capabilities that areengineered to the orthopedic application. Bone graft, whether harvestedfrom the patient or taken from the bone bank, has varying, often unknowndegrees of mechanical properties and overall quality. Similarly, thebone must incorporate into the surrounding bone for long-term clinicalsuccess. If the bone dies or does not generate new bone, the fatiguecharacteristics will be poor and can lead to collapse, loosening, pain,and re-operation. The preferred tantalum material is highly fatigueresistant and maintains its strength for the duration of clinical usage.The mechanical properties should not degrade with time. Since thestiffness properties of the preferred tantalum material are similar tobone, the load pattern to the surrounding bone should be maintainedwithout a compromise of quality.

The preferred tantalum material has a volumetric porosity greater thantraditional prosthetic materials and bone fixation surface coatings.This high porosity allows a more normal restoration of the bone incontact with the porous material, unlike the bone density changephenomenon seen with minimally porous or non-porous implant materials.The solid metals used in current implants are at least ten times stifferthan bone, whereas the tantalum material preferably has a stiffnesssimilar to that of bone.

Initial stability is equally important and is necessary for proper bonein-growth. The tantalum material will preferably have high frictionalcharacteristics when contacting bone. In the early post-operativeperiod, these frictional and structural properties allow the implantdevice to remain very stable.

For soft tissue applications, the properties of porous tantalum have animportant role. Similar to bone, the overwhelming volumetric porosityallows fast penetration of precursor cells and relatively fast formationof soft tissue fibral strands and blood supply. Unlike solid metalscrews, washers or synthetic sutures, porous tantalum achieves theprimary mode of tissue attachment to the implant device while thetissues heal at their own variable pace. The struts of the poroustantalum material interlock with the tissue, offering immediate, secureand functional mechanical attachment. This allows for the necessaryhealing and reproducible tissue incorporation into the porous matrix.The use of a porous tantalum soft tissue anchoring device may thereforeresult in both soft tissue in-growth and bone in-growth for long-termfixation.

One method for forming a base 102 of porous tantalum is described inU.S. Pat. No. 5,282,861 to Kaplan, issued Feb. 1, 1994, and which isherein incorporated by reference. According to the method, the metal,such as tantalum, is deposited on a carbon foam substrate. A reactionchamber encloses a chlorination chamber and a hot wall furnace. Aresistance heater surrounds the chlorination chamber and an inductionheating coil surrounds the reaction chamber to heat the hot wallfurnace. Tantalum metal is located within the chlorination chamber and acarbon foam substrate is positioned within the hot wall furnace.Chlorine gas is injected into the chlorination chamber to react with thetantalum to form tantalum chloride. The tantalum chloride mixes withhydrogen injected into the chamber and then passes through an opening inthe hot wall furnace. The mixture is heated within the hot wall furnaceof a temperature of approximately 1100° C. to produce the followingreacting surface TaCl₅+5/2 H₂→Ta+5 HCl. The surface reaction depositsthe tantalum on the carbon foam substrate to produce a uniform thin filmover the individual ligaments of the substrate. The hydrogen chloride isthen exhausted.

It should be appreciated that although the substrate has been indicatedto be carbon, other carboneous materials, such as graphite, may be used.In addition, other open cell materials, such as high temperatureceramics, may also be used. Also, other layers may be deposited on thesubstrate, such as intermediate layers to provide additional strength.Other aspects of the invention could be the incorporation of a core ofsolid material, such as tantalum or niobium or alloys of each, with theporous substrate fitted around the solid core and with the subsequentdeposition of metal not only covering the substrate but also locking theporous substrate to the solid core.

The base 102 may also comprise porous tantalum formed on a substratematerial. A method for forming the base 102 of porous tantalum on asubstrate material is disclosed in U.S. Pat. No. 6,063,442 to Cohen etal, issued May 16, 2000, and which is herein incorporated by reference.

In another method of forming the base 102, spherical beads or particles(not shown) of Ti or Ti alloy can be charged into a mold or form. Thebeads are preferably of relatively uniform shape. It is within the skillof one in the art to select a bead size range to result in a desiredporous matrix with the desired pore size. The beads can then be exposedto high temperature in a Hot Isostatic Pressing (HIP) process to sinterthe beads into the desired solid matrix form.

The HIP process is carried out in an oven that includes an airlock. Thebase 102 is prepared as described above and placed within the oven,which is then evacuated and charged with an inert (e.g., argon)atmosphere. The oven is heated to the desired temperature while theatmosphere therein is pressurized to the desired pressure. The HIPprocess applies an isostatic pressure through the inert gas (e.g.,argon). By applying sufficient pressure during the heating step, thebeads are fused together at temperature below that which would adverselyaffect the microstructure of the material.

With continued reference to FIG. 1, the hip 100 also includes a ball 106which has a relatively dense, hard and wear resistant outer surfaceregion 108 due to the unique processing and material describedhereinbelow. The ball 106 fits within a prosthetic acetabular socket cup(not shown) and the outer surface region 108 of the ball 106 forms abearing surface with the inner surface of the socket cup. While theporous material, such as porous Ti or Ta making up the base 102 (andball 106) has advantageous bone fusion properties, it would not haveoptimal wear properties for surfaces such as the bearing surface of theball 106.

With reference to FIG. 2, the outer surface region 108 of the ball 106of the hip 100 can be seen in more detail. The outer surface region 108includes a corrosion barrier layer 110 over which a hard dense outermaterial 112 such as Co—Cr is formed.

The outer surface region 108, including the corrosion barrier layer 110and the outer material 112, can be constructed as laser based metaldeposition (LBMD) layers. An example of a LBMD process is LaserEngineered Net Shaping (LENS™), Sandia Corporation of Albuquerque, N.Mex., is described in U.S. Pat. No. 6,046,426 to Jeantette, et al.,issued on Apr. 4, 2000, and which is incorporated herein by reference.Initially, a layer is deposited directly on the ball 106. Thereafter,subsequent layers can be deposited on previous layers in a controlledmanner until a desired surface shape is formed. The material can bedeposited for example as a powdered metal emitted from one or morenozzles. Alternatively, the material could be provided as a wire or as afoil, held in proximity to the base and heated with the laser.

FIGS. 3A-B illustrate the construction of the outer surface region 108of the ball 106 according to a preferred LBMD process. As shown, thecorrosion barrier layer 110 is formed first by depositing a layer ofcorrosion-resistant material 118 such as Ti or Ti alloy onto the ball106, and immediately heating the material with a high power laser 113.Then the outer layer 112 is formed on the corrosion barrier layer 110,again by deposition and laser heating. More detail about a preferredprocess is provided below.

As shown in FIG. 3A, a powdered material feeder (not shown) provides auniform and continuous flow of a measured amount of powdered material118 to the delivery system, or nozzle 114 The delivery system directsthe powdered material 118 toward the ball 106 and directs the powderedmaterial 118 to flow in a converging, conical pattern whereby the apexof such converging, conical pattern intersects the minimum diameter of afocused laser beam (i.e. focus or focal plane) produced by a laser 113such as an Nd YAG laser, all of which is in close proximity to thesurface of the base 102. This generates a melt zone 116, wherein asubstantial portion of the powdered material 118 melts and is depositedon the surface of the ball 106. Those skilled in the art will appreciatethat such powdered material can melt either in flight or upon injectioninto a molten puddle of powdered material. By causing the ball 106 tomove relative to the delivery system or by moving the delivery systemrelative to the ball 106, layers of molten deposited material can bedeposited to form a net-shaped surface.

The deposited corrosion barrier layer 110 may be deposited as a singlelayer, or as multiple layers applied by successive passes of LBMDdeposition. For instance, laminates of corrosion-resistant material(e.g., Ti and/or Ti alloys, etc.) can be formed to create the corrosionbarrier layer 110.

Referring to FIG. 3B, the layer of outer material 112 is formed on thecorrosion barrier layer 110 by a LBMD process as set forth above, thistime using biocompatible material 120 that has a high wear resistance,such as Co—Cr alloy. Again, laminates of high wear resistance materialcan be formed. FIG. 3C is a micrograph at 5× magnification that showsthree layers of Co—Cr alloy 140 deposited by the LBMD process on a bulkmaterial of wrought Co—Cr 142. FIG. 3D is a micrograph at 5×magnification of nine layers of Co—Cr alloy deposited by LBMD on a bulkmaterial of wrought Co—Cr. FIG. 3E is a micrograph at 50× magnificationshowing the bulk wrought Co—Cr alloy. FIG. 3F is a micrograph at 50×magnification showing the LBMD deposited Co—Cr alloy, particularlyshowing the finer grain structure associated with a rapidly cooled LBMDdeposited material.

Either of the layers 110, 112 can also be formed to have a gradient ofmaterial qualities; for example the outer material 112 could be formedto become progressively harder toward the outer surface of the outermaterial 112.

Additional layers can also be added above, below, or between thecorrosion barrier layer 110 and layer of outer material 112 per thedesires of the manufacturer or need in the industry.

The LBMD deposition process is preferably performed in a controlledatmosphere chamber (not shown) which contains an inert gas to inhibitthe formation of surface oxide in the deposition area. This reduces theamount of laser energy needed to achieve full melting of the powder.Although deposition can be performed outside the controlled atmospherechamber, the inert atmosphere will promote full density in the depositedstructure and ultimately lead to improved strength of the appliedsurface material.

It should be appreciated that the laser heats the LBMD depositedmaterial in a very localized manner and for a very short duration.Because of this the heat does not appreciably heat the base material,and thus the heat does not adversely affect the structure of the basematerial. Furthermore, the large heat sink of the ball 106 combined withthe very small area of localized heating causes the heated depositedmaterial to very rapidly cool. This results in a finer grain structurethan would occur with a slower cooling, and also results in carbideinterspersions when conducted in a carbon-rich environment. As thoseskilled in the art will appreciate, fine grain structure and thepresence of carbide interspersions both contribute to improved hardnessand therefore improved wear properties.

In addition, because of the rapid rate of heating and cooling, theapplied material does not tend to excessively flow into porous material,thereby maintaining the desirable porous properties of the porous bulkportion of the device and a relatively small bonding zone between theporous material and the LBMD deposited material. This allows for a thinlayer of LBMD deposited material to be deposited onto porous material.Because this layer of deposited material is thin, implants can befabricated that are optimized in size to limit the amount of bone thatmust be removed to facilitate the bulk of the implant. For example, a 5millimeter thick sheet-like implant with a 3 millimeter thick porousbone ingrowth underside, a 0.5 millimeter bonding zone, and 1.5millimeter bearing surface made from a first layer of Titanium and asecond layer of Cobalt-Chrome can be placed as bearing pads on theproximal tibial plateau as a tibial hemiplasty implant in the knee. Thisconstruct of the 5 millimeter thick implant is significantlybone-conserving compared to traditional 9 millimeter to 20 millimeterthick tibial implants that are currently used to resurface the proximaltibia of the knee.

As mentioned above, the deposited layers may be deposited as multiplelayers applied by successive passes of LBMD deposition. It should bepointed out the heat used to apply each layer and/or the materialcomposition can be adjusted with each pass to achieve a gradient ofmaterial properties if desired. For example, the layer could be appliedso that the applied layers are progressively harder toward the surfaceof the structure.

Another preferred embodiment includes a multi-layer “sandwich” of Co—Cralloy (outer material 112) on titanium (corrosion barrier layer 110) ona porous tantalum or titanium base material. LBMD is used to directlydeposit titanium onto porous tantalum or titanium and Co—Cr onto thepreviously deposited titanium. Illustrative dimensions of such anembodiment follow. The thickness of the porous tantalum can be about0.040 to 1.000 inches, the thickness of the mixed titanium and tantalumlayer can be between about 0.010 and 0.050 inch. The thickness of thetitanium layer can be between about 0.010 and 0.050 inch. The thicknessof the mixed titanium and Co—Cr layer can be about 0.001 to 0.010 inch.The thickness of the Co—Cr layer can be about 0.010 to 0.500 inch. Thus,a sandwich of tantalum, titanium, Co—Cr could range from about 0.071inches to 1.61 inches. Of course these dimensions are provided by way ofexample, and will vary depending on the type and use of the implantdevice.

According to another preferred embodiment, multi-layer structures suchas that described in the preceding paragraph can be formed for couplingto another device such as a commercially available implant. Forinstance, such multi-layer structures can be fusion or diffusion bondedto implants that are made by traditional methods. Thus, for example, theCo—Cr surface of a 0.200 inch three layer structure could be diffusionbonded to a hip or knee implant, as shown in FIG. 10. The porous surfacewould then advantageously be available for coupling to bone of a hostpatient.

In fusion bonding, the substrates are first forced into intimate contactby applying a high contact force. The substrates are then placed in afurnace and annealed at high temperature, after which a solid bond isformed between the substrates. In diffusion bonding, the substrates areforced into intimate contact under high contact force, and heated at atemperature below the melting point of the substrate materials. Fusionbonds involve the complete melting and mixing of both metals. Diffusionbonding can be viewed as a form of fusion bonding but with much lessmelting and mixing of both metals.

With reference to FIG. 4, according to another embodiment of theinvention, the present invention could be used to provide improvedbonding of a first portion 400 of a prosthetic device 402 to a secondportion 404 of the device 402. For example, the first portion 400 mightbe constructed primarily of hard, dense material such as Co—Cr, whilethe second portion 404 might be constructed of a porous material such asporous Ti. Heretofore, bonding of porous Ti with a material such asCo—Cr has achieved poor results. In addition, bonding porous Ti withCo—Cr, resulted in galvanic corrosion across the two dissimilar metals.

According to the present invention, a corrosion barrier layer 406 can bedeposited onto the first portion 400 by laser based metal deposition(LBMD). Thereafter, a layer of Co—Cr 408 can be deposited onto thecorrosion barrier layer, again by LBMD deposition. Co—Cr can be bondedvery well with Co—Cr. Therefore, the LBMD deposited Co—Cr outer surface408 of the second portion 404 can achieve excellent bonding with theCo—Cr of the first portion 400 without any corrosion problems.

FIG. 5 illustrates by way of example and not limitation, various otherpossible devices in which the present invention might be embodied.Devices shown in FIG. 5 include a TMJ joint 500 in situ, an implant forthe great toe 502 (also generally representative of knee, wrist andspinal implants), a dental implant 504 in situ, articulating fingerimplants 506, thumb implants 508, a wrist implant 510 in situ, dentalimplants 512 in situ, a dental implant 514 in situ, a knee implant 516,and a shoulder implant 518 in situ. More detail about each of theseimplants is set forth below.

FIG. 6 is a partial cross sectional view of the toe implant 502 of FIG.5 taken along line 6-6 of FIG. 5. As shown, the implant 502 has a shank600 and a knuckle portion 602 formed from a unitary body of porousmaterial such as tantalum. The porous shank 600 remains exposed forfusion with bone. However, because the knuckle portion 602 is designedto engage a corresponding knuckle of bone, metal or ceramic, the knuckleportion 602 has a smooth outer surface that must be resistant to wear.Using the LBMD process described above, a corrosion resistant layer 604of corrosion-resistant material (e.g., Ti) is formed on at least aportion of the knuckle portion. An outer layer 606 of a wear resistantmaterial (e.g., Co—Cr alloy) is formed over the corrosion resistantlayer 604.

FIG. 7 is a partial cross sectional view of the dental implant 504 ofFIG. 5 taken along line 7-7 of FIG. 5. As shown, the implant 504 has ashank 700 and a tooth coupling portion 702 formed from a unitary body ofporous material such as tantalum. The porous shank 700 remains exposedfor fusion with the jaw bone. However, because the tooth couplingportion 702 is designed to engage an artificial tooth, the toothcoupling portion 702 must be resistant to wear created by the stressesof chewing food. Using the LBMD process described above, a corrosionresistant layer 704 of corrosion-resistant material (e.g., Ti) is formedon at least a portion of the tooth coupling portion 702. An outer layer706 of a wear resistant material (e.g., Co—Cr alloy) is formed over thecorrosion resistant layer 704.

Note that an implant similar to the implant 504 of FIG. 7 can be usedwith the TMJ joint 500 of FIG. 5 to secure the TMJ joint to the jaw andcranium of the host patient. In that case, the implant would be formedof a unitary body of porous material for fusion with bone, the portionof the implant engaging the hinged members would have the corrosionresistant layer and durable outer layer formed thereon by the LBMDprocess. The durable outer layer would resist wear between the implantand the hinged member caused by the stresses of chewing.

FIG. 8 is a partial cross sectional view of one articulating implant 506of FIG. 5 taken along line 8-8 of FIG. 5. As shown, the implant 506 hasa shank 800 and a ball portion 802 formed from a unitary body of porousmaterial such as tantalum. The porous shank 800 remains exposed forfusion with the finger bone. However, because the ball portion 802 isdesigned to engage a corresponding metal socket, the ball portion 802must be resistant to wear. Using the LBMD process described above, acorrosion resistant layer 804 of corrosion-resistant material (e.g., Ti)is formed on at least a portion of the ball portion 802. An outer layer806 of a wear resistant material (e.g., Co—Cr alloy) is formed over thecorrosion resistant layer 804.

FIG. 9 is a partial cross sectional view of the thumb implant 508 ofFIG. 5 taken along line 9-9 of FIG. 5. As shown, the implant 508 has ashank 900 and a knuckle portion 902. Here, the shank 900 is formed ofhydroxy apatite. The knuckle portion 902 is made of metal coupled to theshank 900. The porous shank 900 remains exposed for fusion with bone.However, because the knuckle portion 902 is designed to engage acorresponding knuckle 903, the knuckle portion 902 has a smooth outersurface that must be resistant to wear. Using the LBMD process describedabove, a corrosion resistant layer 904 of corrosion-resistant material(e.g., Ti) is formed on at least a portion of the knuckle portion. Anouter layer 906 of a wear resistant material (e.g., Co—Cr alloy) isformed over the corrosion resistant layer 904.

FIG. 10 depicts the knee implant 516 of FIG. 5. In this embodiment, amulti-layer structure 1000 is independently formed for insertion in thedepression 1002 of the implant 516. The multi-layer structure 1000 isformed of a first layer 1004 of Co—Cr, a middle layer 1006 of corrosionresistant material (e.g., Ti), and an outer layer 1008 of a porousmaterial (e.g., Ta). The multi-layer structure can be fusion ordiffusion bonded to the implant 516 that has been made by traditionalmethods. For example, the Co—Cr surface 1004 of a 0.200 inch three layerstructure can be diffusion bonded to the implant 516. The porous surfaceof the outer layer 1008 is then advantageously available for coupling tobone of a host patient. A description of how to form such multi-layerstructures and how to couple them to implants has been provided above.

Another embodiment of the invention is the deposition of an articulatingbearing material in the form of a composite coating in multiplefunctionally graded layers, in order to enhance the coating performance.FIG. 11 illustrates deposition of a blended material 150 in multiplelayers onto a metal base structure 170. In this example, a Ti/TiCcomposite coating containing 60% TiC (by volume) is deposited on apre-manufactured Ti-6Al-4V implant material. A powder material blend 150of a first material 152 comprising 40 Vol. % CP Ti (+100/−325 mesh) anda second material 154 comprising 60% TiC (+100/−325 mesh) is prepared bythoroughly mixing the constituent powders using ball milling or similarpowder mixing techniques. The deposition process begins with directing afocused Nd-YAG laser beam 113 onto a substrate 170 placed on the buildplatform capable of computer-controlled motion. The computer (not shown)utilizes a deposition file which specifies optimal deposition parametersalong with required coating thickness and deposition area profiles.Deposition parameters include: laser power, scan speed, scan spacing,scan orientation, layer thickness, powder feed rate, and others.Required coating thickness includes desired final coating thickness plusany machining allowance.

The laser 113 generates a small molten pool (typically 0.25-1 mm indiameter and 0.1-0.5 mm in depth) 116 on the substrate 170. Preciseamounts of the powder material blend 150 are injected directly throughnozzles 114 into the melt pool 116 using a powder feeder (not shown).The molten pool 116 solidifies rapidly as the laser beam 113 moves away,forming a thin track of solidified metal welded to the material belowalong the line of laser scanning. A layer 156 of coating is generated bya number of consecutive overlapping tracks. A second layer 158 isgenerated in the same way, by a series of consecutive overlappingtracks, as are additional layers 160, 162. After each layer is formed,the laser head 113, along with the powder delivery nozzle 114, movesupward by one layer thickness and the subsequent layer is generated.This process is repeated until a coating, composed of multiplefunctionally graded layers, of desired thickness is deposited. In thisexample, four layers are deposited; however any number of layers fromone on up can be deposited, depending on properties desired for the useof the implant. The deposition process occurs inside an enclosed chamberfilled with argon to prevent oxidation of the liquid metal. The partbuilding process is fully automatic and can run unattended.

FIG. 12 illustrates a variation on the above process, in which thearticulating bearing powder material blend 150 is deposited in a singlelayer 156 upon the base structure 170. This technique could be used whena very thin layer of coating is desired, depending on the type and useof the implant. FIG. 13 illustrates a variation in which a firstarticulating bearing material 166 instead of the blend is used, but isdeposited in a series of layers 156, 158, 160, 162, forming layers ofthe articulating bearing material on the base structure 170.

Referring to FIG. 14, a combination of the first material 152 and thesecond material 154 are deposited in multiple layers upon the basestructure 170. One nozzle 114 delivers the first material 152, such asTitanium, as the other nozzle 114 delivers the second material 154, suchas Titanium carbide. As each successive layer 156, 158, 160, 162 isdeposited, the computer controls the distribution of the first andsecond materials 152, 154, varying them by layer to produce functionallygraded layers.

In the embodiments illustrated in FIGS. 11, 12, 13 and 14 thearticulating bearing material is deposited on a base structure 170. Thebase structure 170 could be, but is not limited to, a hip implant, a TMJjoint implant, an implant for the great toe, a knee implant, a wristimplant, spinal implants, a dental implant, an articulating fingerimplant, a thumb implant, and a shoulder implant. In addition, theprocess used to deposit the articulating bearing material can include,but is not limited to any energy source that can provide sufficientenergy locally to melt and deposit the additive material, such as LBMD,plasma, or electron beam energy. A welding based cladding process mayalso be used. Also, the first and second materials, their particle sizesand distributions may be varied to produce different microstructuralcoatings, depending upon the desired coating characteristics for thespecific application.

Another embodiment of the invention includes the deposition of anantimicrobial material in combination with the metal base material, ontothe implant base structure. One antimicrobial material may be silver inthe form of elemental silver, a silver salt or a silver intermetallic.The antimicrobial properties of silver have been well documented, withnumerous studies measuring the efficacy of silver in the reduction ofbacteria including Escherichia coli, Staphylococcus aureus, andStaphylococcus epidermis, among others. However, other suitableantimicrobial materials may be used in place of silver, including butnot limited to gold, platinum, palladium, iridium, copper, tin,antimony, bismuth, zinc, salts thereof, and intermetallics thereof.

FIG. 15 illustrates the deposition of a combination of a metal and anantimicrobial material onto the surface of an implant. A blend 180 of afirst metal 182 and an antimicrobial material 184 is directed in streamsfrom the nozzles 114 to the surface of the base structure 170. The firstmetal 182 may be the same metal which forms the base structure 170, andmay comprise cobalt-chrome, tantalum, titanium, platinum, zirconium,niobium, stainless steel, or an alloy thereof. The antimicrobialmaterial 184 may be elemental silver, a silver salt, a silverintermetallic, or any other antimicrobial metal or material which issuitable to be combined with the first metal 182. The laser 113 meltsthe blend 180, creating the melt pool 116. The molten pool 116solidifies rapidly as the laser beam 113 moves away, forming a thintrack of solidified blend 180 welded to the material below along theline of laser scanning. A layer 200 of the deposit is generated by anumber of consecutive overlapping tracks. A second layer 202 isgenerated in the same way, by a series of consecutive overlappingtracks, as are additional layers 204, 206. After each layer is formed,the laser head 113, along with the powder delivery nozzles 114, moveupward by one layer thickness and the subsequent layer is generated.When all layers have been deposited, a multi-layered deposit 190 remainson the surface of the base structure 170.

FIG. 16 illustrates a variation on the above process, in which theantimicrobial material blend 180 is deposited in a single layer 208 uponthe base structure 170. This technique could be used when a very thindeposit is desired, depending on the type and use of the implant. Adeposit as thin as 25 microns may be created by this process. Postdeposition processes which include but are not limited to annealing,etching in acid or base, or oxygen plasma can be used to control thelayer thickness below 25 microns, as well as control surface porosity.

The antimicrobial material 184 and first metal 182 may also be depositedin functionally graded layers. Referring to FIG. 17, a combination ofthe first metal 182 and the antimicrobial material 184 are deposited inmultiple layers upon the base structure 170. One nozzle 114 delivers thefirst metal 182, such as titanium, as the other nozzle 114 delivers theantimicrobial material 184, such as elemental silver. As each successivelayer 210, 212, 214, and 216 is deposited, the computer controls thedistribution of the first metal 182 and antimicrobial material 184,varying them by layer to produce functionally graded layers.

Such functionally graded layers can be used to accomplish a number ofdifferent results. According to one example, the concentration of theantimicrobial material 184 may gradually increase as new layers areadded so that the outermost layers possess the most potent antimicrobialproperties. According to another example, the harder of two materialsbeing combined may be deposited in greater concentrations as successivelayers are formed to provide an increased hardness of the outermostlayers. Such a structure may be particularly useful for the formation ofarticular surfaces. According to another example, the more porous of twomaterials being combined may be deposited in greater concentrations assuccessive layers are formed to provide an increased porosity of theoutermost layers. Such a structure may be particularly useful for theformation of bone in-growth surfaces.

Combinations of the above-described examples are also possible.Additionally, more than two different materials may be combined; suchmaterials may be deposited in functionally graded layers that accomplishmultiple desirable property changes along the thickness of the coating.Such functional gradients may be employed regardless of whether any ofthe materials being deposited has antimicrobial properties.

FIG. 18 illustrates another embodiment of the invention, in which aplurality of zone-specific deposits comprising different compositions ismade on the same base structure, in this example a knee implant. Thedepression 1002 of the implant 516 is shaped to hold the multi-layerstructure 1000, which comprises a deposited outer layer 1008 of a porousbone ingrowth material. On an anterior curved surface 1010 of theimplant 516, a deposited outer layer 1012 comprising a bearing surfacematerial such as titanium carbide is formed by the LBMD process. Onlateral edges 1014 on either lateral side of the implant 516, adeposited outer layer 1016 comprising an antimicrobial material isformed by the LBMD process. The lateral edges 1014 are not designed forarticulation or bone apposition, but may instead help to reduce thelikelihood of infection of the surrounding tissues due to theantimicrobial of the deposited outer layer 1016.

In other examples, antimicrobial material may be combined with eitherthe porous bone ingrowth material, or the bearing surface material, orboth, and deposited on the implant. It is appreciated that variouscombinations of base structure material, porous bone ingrowth material,bearing surface material, and antimicrobial material may be made anddeposited on the implant, depending on the desired characteristics forthe particular zone of the implant surface, be it bone ingrowth,bearing, antimicrobial, or a combination thereof. Furthermore, enhancinghardness, porosity, or antimicrobial properties are not the onlyfunctions that may be served by biomedical implant coatings formed viaLBMD; indeed, such coatings may help to enhance biocompatibility of theimplant, provide a desired level of radio-opacity, or even providefunctional geometrical shapes that may be used for interconnection withother implants, bone anchorage, or other purposes.

While the present invention has been disclosed in its preferred form,the specific embodiments thereof as disclosed and illustrated herein arenot to be considered in a limiting sense, as numerous variations arepossible. The invention may be embodied in other specific forms withoutdeparting from its spirit or essential characteristics. No singlefeature, function, element or property of the disclosed embodiments isessential. The scope of the invention is, therefore, indicated by theappended claims rather than by the foregoing description.

1. A biomedical implant comprising: a metal base structure; and adeposit formed onto the metal base structure in functionally gradedlayers, the deposit comprising an antimicrobial material and a firstmetal, wherein the outermost layers possess the most potentantimicrobial properties, wherein depositing the deposit is accomplishedin an inert gas atmosphere.
 2. A biomedical implant as in claim 1,wherein the antimicrobial material comprises at least one of elementalsilver, gold, platinum, palladium, iridium, copper, tin, antimony,bismuth, zinc, salts thereof, or intermetallics thereof.
 3. A biomedicalimplant as in claim 1, wherein the first metal comprises at least one ofcobalt-chrome, tantalum, titanium, platinum, zirconium, niobium,stainless steel, or alloys thereof.
 4. A biomedical implant as in claim1, wherein the deposit has a thickness greater than 25 microns.
 5. Abiomedical implant as in claim 1, wherein the deposit is furthermodified by a process selected from the group consisting of annealing,etching in acid, etching in base, and oxygen plasma.
 6. A method forconstructing a biomedical implant, the method comprising: a metal basestructure; and a deposit formed onto the metal base structure by LaserBased Metal Deposition (LBMD) in an inert gas atmosphere, wherein thedeposit comprises an antimicrobial material and a first metal, whereinthe deposit comprises functionally graded layers, the outermost layerspossessing the most potent antimicrobial properties.
 7. A biomedicalimplant as in claim 6, wherein the antimicrobial material comprises atleast one of elemental silver, gold, platinum, palladium, iridium,copper, tin, antimony, bismuth, zinc, salts thereof, or intermetallicsthereof.
 8. A biomedical implant as in claim 6, wherein the first metalcomprises at least one of cobalt-chrome, tantalum, titanium, platinum,zirconium, niobium, stainless steel, or alloys thereof.
 9. A biomedicalimplant as in claim 6, wherein the metal base structure comprises a baseshaped to be secured to a body part of a patient.
 10. A biomedicalimplant as in claim 6, wherein the deposit has a thickness greater than25 microns.
 11. A biomedical implant as in claim 6, wherein the depositis further modified by a process selected from the group consisting ofannealing, etching in acid, etching in base, and oxygen plasma.
 12. Abiomedical implant comprising: a base structure formed from abiocompatible metal material; and at least one deposit formed onto thebase structure by Laser Based Metal Deposition (LBMD) in an inert gasatmosphere and in functionally graded layers, wherein the outermostlayers possess the most potent antimicrobial properties, wherein thedeposit includes at least one selection from the group consisting of afirst metal, an antimicrobial material, a porous bone ingrowth material,a bearing material, and combinations thereof.
 13. A biomedical implantas in claim 12, wherein the antimicrobial material comprises at leastone of elemental silver, gold, platinum, palladium, iridium, copper,tin, antimony, bismuth, zinc, salts thereof, or intermetallics thereof.14. A biomedical implant as in claim 12, wherein the first metalcomprises at least one of cobalt-chrome, tantalum, titanium, platinum,zirconium, niobium, stainless steel, or alloys thereof.
 15. A biomedicalimplant as in claim 12, wherein the first metal comprises at least oneof cobalt-chrome, tantalum, titanium, platinum, zirconium, niobium,stainless steel, or alloys thereof.
 16. A biomedical implant as in claim12, wherein the bearing material comprises a biocompatible compositioncomprising cobalt and chromium.
 17. A biomedical implant as in claim 12,wherein the metal base structure comprises a base shaped to be securedto a body part of a patient.
 18. A biomedical implant as in claim 12,wherein each deposit has a thickness greater than 25 microns.
 19. Abiomedical implant as in claim 12, wherein at least one deposit ismodified by a process selected from the group consisting of annealing,etching in acid, etching in base, and oxygen plasma.